Medical endoprostheses or implants for a wide variety of applications are known in a great variety from the prior art. Endoprostheses in the sense of the present invention include, for example, endovascular prostheses, e.g., stents, fastening elements for bones, e.g., screws, plates or nails, surgical suture materials, intestinal clamps, vascular clips, prostheses for use in the area of hard and soft tissue as well as anchoring elements for electrodes, in particular pacemakers or defibrillators.
Stents are used especially commonly today as implants for treatment of stenoses (vascular occlusions). They have a basic mesh which is tubular or hollow cylindrical and is open at both longitudinal ends. The tubular basic mesh of such an endoprosthesis is inserted into the vessel to be treated and serves to support the vessel. Stents have become established for treatment of vascular diseases in particular. Through the use of stents, constricted areas in the vessels can be dilated, resulting in a larger lumen. Although an optimum vascular cross section can be achieved through the use of stents or other implants, and this is one of the primary requirements for successful treatment, the permanent presence of such a foreign body initiates a cascade of microbiological processes which can lead to a gradual overgrowth of the stent and, in the worst case, to a vascular occlusion. One starting point toward solving this problem is to manufacture the stent and/or other implants from a biodegradable material.
The term “biodegradation” is understood to refer to hydrolytic, enzymatic and other metabolic degradation processes in a viable organism, caused primarily by the body fluids coming in to contact with the biodegradable material and leading to gradual dissolution of the structures of the implant containing the biodegradable material. Through this process, the implant loses its mechanical integrity at a certain point in time. The term “biocorrosion” is often used as synonymous with the term biodegradation. The term “bioresorption” includes the subsequent absorption of the degradation products by the living organism.
Materials suitable for the basic mesh of biodegradable implants may contain polymers or metals, for example. The basic mesh may comprise several of these materials. The feature these materials have in common is their biodegradability. Examples of suitable polymeric compounds include polymers from the group of cellulose, collagen, albumin, casein, polysaccharides (PSAC), polylactide (PLA), poly-L-lactide (PLLA), polyglycol (PGA), poly-D,L-lactide-co-glycolide (PDLLA-PGA), polyhydroxybutyric acid (PHB), polyhydroxyvaleric acid (PHV), polyalkyl carbonates, polyorthoesters, polyethylene terephthalate (PET), polymalonic acid (PML), polyanhydrides, polyphosphazenes, polyamino acids and their copolymers as well as hyaluronic acid. Depending on the desired properties, the polymers may be in pure form, in derivatized form, in the form of blends or copolymers. Metallic biodegradable materials are based primarily on alloys of magnesium and iron. The present invention preferably relates to implants whose biodegradable material contains at least partially a metal, preferably magnesium or a magnesium alloy.
Stents having coatings with different functions are already known. Such coatings serve to release medicines, to arrange an X-ray marker or to protect the underlying structures, for example.
In the implementation of biodegradable implants, the degradability should be controlled in accordance with the desired treatment and/or use of the respective implant (coronary, intracranial, renal, etc.). For many therapeutic applications, for example, an important target corridor is for the implant to lose its integrity over a period of four weeks to six months, for example. Integrity here is understood to be mechanical integrity, i.e., the property whereby the implant suffers hardly any mechanical losses at all in comparison with the undegraded implant. This means that the implant still has such great mechanical stability that the collapse pressure, for example, drops only slightly, i.e., to at most 80% of the nominal value. The implant may thus still fulfill its main function, i.e., keeping the blood vessel open, if it still has integrity. Alternatively, the integrity may be defined by the fact that the implant has such great mechanical stability that, in its stressed state in a blood vessel, it is hardly subject to any geometric changes, e.g., it does not undergo any mentionable collapse, i.e., it still has at least 80% of the dilatation diameter under load or, in the case of a stent, hardly any of the struts are broken.
Biodegradable magnesium implants, in particular magnesium stents have proven to be especially promising for the aforementioned target corridor of degradation, but, first of all, they lose their mechanical integrity, i.e., the supporting effect, too soon and on the other hand have a greatly fluctuating loss of integrity in vitro and in vivo. This means that in the case of the magnesium stents, the collapse pressure drops too rapidly over time and/or the reduction in the collapse pressure has too much variability and therefore cannot be determined.
Various mechanisms for controlling the degradation of magnesium implants have already been described in the prior art. These are based, for example, on organic and inorganic protective layers or combinations thereof which counteract the human corrosion medium and present resistance to the corrosion processes taking place there. Approaches known in the past have been characterized in that they achieve barrier layer effects which are based on a spatial separation that is as complete as possible between the corrosion medium and the metallic material, in particular the metallic magnesium. The degradation protection is thus ensured by protective layers having various compositions and by defined geometric distances (diffusion barriers) between the corrosion medium and the magnesium base material. Other approaches are based on the alloy components of the biodegradable material of the implant body which influence the corrosion process by displacing the location in the electrochemical voltage series. Other approaches in the field of controlled degradation induce predetermined breaking effects by applying physical changes (e.g., local changes in cross section) and/or chemical changes in the stent surface (e.g., multiple layers having different chemical compositions). However, with the approaches mentioned so far, it is usually impossible to have the dissolution that occurs due to the degradation process and to have the resulting breakage of webs occur within the required time frame. The result is that either the onset of degradation is too early or too late and/or there is too much variability in the degradation of the implant.
Another problem that occurs in conjunction with passivation coatings is based on the fact that stents or other implants usually assume two states, namely a compressed state with a small diameter and an expanded state with a larger diameter. In the compressed state, the implant can be inserted into the vessel to be supported by means of a catheter and can be positioned at the site to be treated. At the site of treatment, the implant is then dilated by means of a balloon catheter and/or (when using a shape memory alloy as the implant material) converted to the expanded state, e.g., by heating to a temperature above the transition temperature. On the basis of this change in diameter, the body of the implant here is exposed to a mechanical stress. Additional mechanical stresses on the implant may occur during production or in the movement of the implant in or with the vessel into which the implant is inserted. With the known coatings, this yields the disadvantage that the coating cracks during deformation of the implant (e.g., forming microcracks) or is even partially removed. This may cause an unspecified local degradation. Furthermore, the onset and rate of degradation depend on the size and distribution of the microcracks, which are formed due to deformation and, as defects, are difficult to control. This leads to a great scattering in the degradation times.
The document DE 10 2006 060 501 discloses a method for manufacturing a corrosion-inhibiting coating on an implant made of a biocorrodible magnesium alloy and an implant obtainable by this method in which, after the implant has been prepared, the implant surface is treated with an aqueous or alcoholic reaction solution containing one or more ions selected from the group comprising K+, Na+, NH4−, Ca2+, Mg2+, Zn2+, Ti4+, Zr4+, Ce3+, Ce4+, PO33−, PO43−, HPO42−, H2PO4−, OH−, BO33−, B4O73−, SiO32−, MnO42−, MnO4−, VO3−, WO42−, MoO42−, TiO32−, Se2−, ZrO32− and NbO4−, where the concentration of ion(s) is in the range of 10−2 mol/L to 2 mol/L. The treatment of the implant surface with the aforementioned reaction solution necessitates anodic oxidation of the implant. It is performed either with or without an external current source (externally currentless). However, the examples of methods and electrolyte compositions described in this document do not meet expectations with regard to degradation behavior and dilatation ability without destruction of the layer when used for a magnesium stent.